Field of the Art
The present invention relates generally to prosthetic legs and, in particular to modular knee joints for prosthetic legs. In particular the present invention relates to a prosthetic leg for those having had a transfemoral amputation, namely those having an above the knee stump, and particularly to a prosthetic leg for transfemoral amputees for whom 50% or more of their femur has been amputated. The present invention also lends itself to prosthetic modularity.
Discussion of the State of the Art
Amputees who use a lower limb prosthesis equipped with a knee joint, find certain movements difficult to perform. For example, the fact that the prosthesis comprises a mechanical apparatus performing the duty of the joint that is substantially distal to the hip can induce a lack of balance. Indeed, patients having a substantial reduction in femoral length will suffer greatly since the prosthetic attachment typically has a hollow region, a socket, in which to insert the remaining length of thigh, the stump; the stump will bear upon the distal aspect of the femur and the femoral lever action will be reduced, in correspondence with the reduction in length. Indeed the pressures can become unbearably high-often resulting in an occurrence of friction burns arising through sliding contact with the prosthesis stump. In fact the length to width ratio of the stump can come close to 1:1 as opposed to a 2:1 or 3:1 as is typical for a normal human femur. It will be understood that the center of mass of a state of the art knee mechanism in a lower limb prosthesis is distal to the same knee center, which allows for a universal application of the prosthetic joint; indeed, any aspect of such a joint proximal to the knee center is kept as short as possible to support a wide as possible commercial and clinical application independent of femoral stump length. Naturally there is a transition across length, but practice bears out that a “heavy” prosthesis is beneficial to the long-femured transfemoral amputee for reasons of body mass symmetry, and very cumbersome and uncomfortable for the short-femured transfemoral amputee, a distinct sub-class of transfemoral amputee the disclosure is made.
In the provision of an above-the-knee prosthesis with locking and/or yield characteristics, one can employ mechanical, pneumatic, or hydraulic devices. There are five distinct categories of prosthetic mechanisms available for those with knee and above knee limitations. The primary design—as in the most common type—comprises a polycentric mechanism wherein the virtual center of rotation is well above the anatomical knee axis—the mechanical apparatus, and its center of mass (approx. 850 gr, 1½ lb.), is distal to the anatomical knee center. Another prosthetic mechanism is the hydraulic swing'n'stance (“SNS”) mechanism developed in the late 1960's by Hans Mauch. This design operates in a stance phase of operation in a default mode; the mechanism is brought into swing on hyperextension of the femur. The Mauch SNS has an internal mechanism where the overstretching of a hydraulic assembly lifts a valve-operating member and enables the leg to operate in a low resistance swing phase. This, again, is enabled at a point distal to the anatomical joint center. A further prosthetic type comprises a weight-activated type of knee joint such as described in GB2464620 (Boender). A fourth type of prosthesis is exemplified by a number of relatively expensive, external battery powered, electromechanical knee joints having sensors and dampers operating under software control for control of the same joint function, albeit with a relatively high mass (approx. 1.4 kg, 3 lb.), distal to the knee center. A fifth type of prosthesis is the fluidic controlled knee joint, wherein by means of fluid flow and mechanical linkages, appropriate negative feedback to the control of movement can be realized, which suffers the same penalty of a relatively high mass, distal to the knee center. It will be readily understood that the more distal and the larger the mass of the knee joint becomes, the effects of unwanted reactive pressures which act on the distal end of a short amputation stump are exacerbated, even at low ambulatory speeds.
The prevalence of prostheses is increasing as people are living longer and in view of the provision of prostheses in respect of previously life threatening injuries and conditions arising from accidents where limb injuries occur are and improvements, for example, in diabetic management, become more widely available. Prostheses manufacturers have adopted certain interoperability standards whereby a modular knee joint from one manufacturer can be replaced with a unit form another manufacturer, meaning that the knee joint is a dismountable and exchangeable part of a whole-leg prosthesis. While each knee joint prosthesis cooperates with other parts of the prosthesis such as the foot in mid-swing, and operable to lower the mass of the user's body down a stair, it is a distinct modular device with a proximal end and a distal end which are the boundaries of the space envelope of the knee joint per se. Naturally a modular joint, as an individual product, has a center of mass and the weight of a modular joint is typically declared in the sales literature and is provided with distributed product. Typical prior art modular knee joints have their center of mass distal to the knee center (or knee axis) about which these are operable. Parts of a prosthesis, such as a foot, shin tube, torque absorbers etc. are referred to as prosthetic components, which are necessary to build a prosthesis, but can be considered separately to the present invention. These components need to be moved throughout space during swing phase of the prosthesis, and the precise movement—with particular reference to the angular relationship between these components relative to the stump attachment device—is controlled by the knee joint. The knee joint has its own center of mass that can be identified in isolation to any housing or exo-skeletal part of the knee joint, which along with the mechanical characteristics of the joint to include aspects such as damping (hydraulic or otherwise), friction (hydraulic or otherwise), position of pivot axis—or axes in the case of a poly-centric joint etc.). For the purposes of scope of disclosure, hip disarticulation prostheses will be understood to count as a special form of short transfemoral amputation prosthesis. Hydraulic knee devices benefit from the liquid being incompressible and can provide friction dependent upon speed; locking and yielding functions can simply be provided by different settings of a variable orifice.
Control of a prosthetic device through the yielding and locking functions of the prosthesis, must have near perfect reliability, particularly on stairs, inclines, and in stumbling situations, in order to prevent accidents. For example, in a weight activated knee joint, a change in posture (or equivalent) is required to apply weight to the joint, whereby to enable the joint to switch from its default free swing mode into a stance mode for weight acceptance, to prevent collapse. Similarly, the default stance joints need to be brought into a low torque or low resistance mode to facilitate the swing phase. Referring to FIGS. 1i and 1ii, there is shown a simplified view of one type of known prosthesis 10, comprising a thigh member 17—complete with insert, also known as a socket 18, a knee 13 and shin 14, in positions of toe-contact but under different circumstance. FIG. 1i shows the leg in stumble mode, where a swing extension movement is interrupted by sudden toe contact with the ground. In an attempt to prevent a fall, the body mass must be shifted to the prosthesis, whereby to allow placement of the sound limb ahead of the body so that the forward momentum of the body will not result in a fall. This figure also shows how the forces pass through the forefoot in stumble mode—which correspond with the forces that pass through the foot in toe-off mode. When a ground reaction force passes slightly posterior to the knee center, then the knee joint will readily bend until movement is arrested. At toe-off, the knee joint will also readily bend as in swinging the leg. In the case of a stumble the knee joint will bend further and collapse, being a most undesirable outcome. Typical state of the art weight activated joints do not have stumble recovery options as in FIG. 1i, even though it is clearly desirable for such prostheses to be provided with a stumble recovery mode. Indeed, neither such a mechanism, nor the polycentric mechanism allow any change in bending resistance once the load is on the forefoot/toe. Accordingly, a stumble will certainly lead to a collapse. Notwithstanding this, the short stumped amputee will be bereft of any intrinsic stumble recovery ability, where a long stumped individual could arrest a fall by employing their effectively stronger hip muscles. FIG. 1ii shows how forefoot 11, having a ground reaction force 12 acting through the knee 13 and which would otherwise be a poor location for engagement of the stance phase, except that it enables the knee to commence swing action. It will be realized that during normal swing function, a stance resistance would be most undesirable.
Indeed the main body of these joints is typically located distal to the knee axis. It will be appreciated that the short stumped, above knee amputee is not well served by these devices of the prior art, especially those that use a default stance mechanism wherein bodyweight on the forefoot at toe-off is required to release the knee into a swing mode. For example, the need for specific stump movement to drive the prosthesis into swing under a residual toe-load can be too much. Instead a weight activated knee joint is generally found to be preferable, where body weight is usefully employed to stabilize the stance mode, and upon simple removal of body weight defaults to a swing mode.
Any prosthesis will have mass. While one important issue is the effective weight arising from gravity, a more troublesome issue arises upon use and, in particular, upon a forward swing of the prosthesis. Unlike the weight of the prosthesis that remains the same irrespective of distribution of mass, the perception of “weight” in kicking the leg forward with the thighbone is very much dependent on the mass distribution. Those skilled in the art will readily confirm that the inertia of a given mass M that is accelerated by a moment, (that is a force applied perpendicularly at a distance r) is given as Mr2, which means that, if a similar mass is twice the distance away from the source that pushes it, the inertia (or the sensation of its heaviness) is four times as great. In the case of a prosthesis, there will always be a foot of some description that must be at a given distance from the hip, to create the right leg length for balance, body symmetry, and an energy efficient gait. Therefore the location of the mass of the foot cannot be altered, and the only way to reduce the inertia is to reduce the mass of the foot and shoe. The situation for the knee joint is quite different; the location of the center of mass of the knee joint can be brought close to the knee center (by selecting a small joint), and by selecting a small mass (small knee joint) at a cost of losing function such as stumble recovery, yielding under body weight in downstairs walking, because in engineering functionality typically adds mass.
Known prosthetic knee joints that are activated only by the application of weight cannot substantially assist the short stumped individual. The weight placed on the toe, as in toe-off, will typically cause disengagement of any stance mode there otherwise might be. This has to be so, because at toe-off, prior to swing, when residual body weight is on the toe, the knee joint must be free to release for swing, and this condition of weight acceptance is, using known devices, indistinguishable from any attempted weight acceptance during attempted stumble recovery with respect to the knee mechanism. In summary, known prosthetic knee & limb devices are either light weight—a benefit, with meager function or safety—a significant disadvantage, but if the engineering requirements are met—a benefit, the inertial mass of the device is too high for comfort—a significant disadvantage, liable to cause unwanted side effects such as the occurrence of friction burns on the distal stump.
Such prior mechanisms, whether mechanically or electronically controlled and operated, rely upon the application of forces by the prosthesis upon the prosthetic apparatus distal to the knee center, noting that the center of mass of the prosthesis is also distal to the knee center. For example, Yuichi (US2005015156) claims to teach of an above-knee prosthesis that permits a user to control knee flexion or extension and to enable voluntary lock and release of the knee joint, at any angle of bend. Indeed, an objective of Yuichi is to provide a natural gait. The prosthesis to Yuichi comprises a thigh frame assembly that receives a thigh stump; a leg frame assembly operably associated with a foot; a hinge interconnects the assemblies to form an artificial knee joint. A closed hydraulic system provides variable resistance to the bending of the artificial knee joint in correspondence with anterior-posterior (AP) movement of the thigh stump. The AP movement of the thigh stump is conveniently arranged to control a flow rate valve by means of a linkage, sliding or screw assembly. The flow rate control valve varies the resistance provided by the closed hydraulic system: pressing the thigh stump backwards within the thigh frame assembly increases resistance and slows knee bending until the knee locks, while pressing the same forwards decreases resistance and allows the artificial knee joint to yield to outside forces, such as gravity and/or stump thrust, the prosthesis can pivot freely about the knee hinge.
Yuichi operates solely on the AP force vectors present in the hip flexion and extension efforts and all embodiments teach of isolation of AP forces from any axial forces, even when the incline of the sensor is slightly tilted, i.e. not perpendicular to the femoral axis: separation occurs between force vector components in the operating plane of the sensor and force vector components in the plane normal to the operating plane. Confusingly, Yuichi defines “AP” in terms of a plane/direction of operation of the sensor device, which makes the discussion about any tilt of this direction relative to the body immaterial. Also confusingly, Yuichi suggests of control by means of hip musculature (¶45), despite that fact that in real gait, for example, upon heel strike with the ground, a force vector naturally passes posterior to the knee center, with reference to Newton's third law. Due to gravity, a body will constantly maintain contact with the ground; the reaction force arising from the ground is the ground reaction force (GRF). The GRF, along with weight, is an important external force. The GRF is normally determined by means of a force-plate, which comprises four tri-axial force sensors that measure the force acting between the foot and the ground in three axes: transverse (Z), anterior-posterior (X), and vertical (Y). The resultant sum of all the reactions from the ground is equivalent to the sum of the four forces measured by the sensors (for more detail, refer to material relating to Biomechanics course 3150 per “BioMedical Engineering OnLine”. The implications of this is that the prosthetic tibia by direction of ground reaction force (GRF) must attempt to move rearwardly with respect to the distal stump, or distal femur slide forward (anteriorly) over the knee apparatus, the GRF vector being shown upon heel touch down of a left leg with reference to FIG. 1iii. Thus, in accordance to Yuichi, would cause the knee joint to be in a low resistance mode and collapse would commence. When the artificial limb apparatus allows commencement of partial collapse, the ground reaction force would start to tilt forward in reaction to a thigh extension reflex, and in accordance to Yuichi's teaching the knee would stabilize and due to the body weight sliding backwards relative to the knee axis, the hydraulic resistance would soon increase to its maximum, but not be under useful voluntary control. Indeed, the AP forces to Yuichi are defined relative to the continually moving reference “vertical”, corresponding to a longitudinal axis of the femur, in alignment with the coronal plane and this is in contrast with any standard medical reference of the term “AP”.
Any claimed voluntary control that is possible with such devices with respect to a level of resistance over knee bending in sitting down does not necessarily become apparent in practice and it is believed that the resistive bending moments will have an “AP” force component equal to the product of the knee bending resist moment and an inverse factor of the distance from hip to knee center, which for a typical male would correspond to the forces arising from half the weight, say 450N (100 lb force). Because the force returned in the sensor device is equal to the pressure the user exerts on it, the user must also provide such levels of force to the sensor device to get a level of resistance. When walking down a stair, the thigh is at approximate 15° relative to vertical, which means that in the time frame of single leg weight bearing, the “AP” force component delivered to the control device is only 25% (sin 15°) of the body weight, meaning that the resistive torque that slows down the descent of the body will permit a g-force of 75%×9.81 to act on the body, which would result in an uncontrolled descent.
In another scenario, where a limb swings through mid-swing, it is known that the bending resistance of the knee must be a maximum at mid-swing, which in terms of such prior devices can only be provided by a posterior force exerted on the knee mechanism in mid-swing to provide this level of knee control consistent with the teachings. This means that the forward swinging thigh must provide a sudden rearward kick in mid-swing to provide that resistance to mid-swing flexion control, which is naturally not possible, or likely to provide an ungainly spectacle. In the alternative there is no mid swing flexion control against excessive knee flexion seen in so many prosthetic devices in mid-swing.
Further, Yuichi employs the AP plane as per AP plane of the knee device as a means of control, using hip extension and hip flexion forces to control the knee joint's operational characteristic. When the ground reaction force, (a mixture of body weight, resisted body momentum when touching the ground, and any voluntary and involuntary muscle actions) passes the knee joint perpendicular to the controlling AP plane, the behavior of the knee device is either indeterminate, or biased by a resilient element, whereby, for example to maintain a level of resistance low unless the resilient element is compressed by hip effort and, thus, make the behavior of the knee device more predictable. However, for individuals with a short transfemoral stump, the precise control with respect to the resilient element is difficult at best, taking into account the characteristics that if the device can be set such that the required force to control is the same as the force returned by the controller, and that individuals with short transfemoral stumps often have limited strength in their leg stump to control a distal device.
With reference to FIG. 1iv, which shows a Pedottii Diagram (commonly referred to as the “butterfly diagram”) and FIG. 1iii, a GRF is initially tilted backwards, more or less in line with a limb. The body is naturally stable if the GRF passes anterior to the knee center, but the “AP” force vector from the GRF, is oppositely directed with respect to the hip's “AP” force vector and in this scenario the “AP” force vector as per Yuichi would force the knee into stability, which would be undesirable. In the event that the GRF vector passes posterior to the knee axis, then the “AP” force vector-set of shin-and-hip would be opposite and would permit the knee to bend, since the knee would react as if the hip was driving the leg forward. This would result in an immediate knee collapse, followed by a rapid transition to a posterior direction of the “AP” force by the hip and the knee would find stability again and a fall would be arrested, but all at a cost of unstable performance.
FIG. 1iv is a graph that illustrates mean anteroposterior (Fx) GRF component waveform exerted by right and left legs (dotted line) in walking at normal speed. It can be determined that the AP force, in early stance, acts in a backward direction so as to provide a braking action on the body. As would be apparent to a skilled man, this would cause the teaching of certain prior art systems (e.g. Yuichi) teaching to create a free knee joint, which is contrary to the desired response. In late stance the body propels itself forward, but in this stage the knee is inherently stable since knee hyperextension is naturally prevented. It is believed that isolation of a single force “AP” component has not been shown to be sufficient for the control of a prosthesis.
In another prior system—per Blatchford (GB2328160A)—a device is disclosed with an axis proximal to the knee center, which suffers similar logical conflicts as with the teaching to Yuichi with a force-sensing element proximal to the knee center to assist in the control of resistive behavior of the control apparatus distal to the knee center. A device like Blatchford relies on force vectors to pass posterior to the auxiliary axis to generate a desired compression force in an associated sensor, which sensor will in turn switch on the resistance in the distal apparatus. While the sensor is proximal to the knee center, the actual controlling apparatus is distal to the knee center, which is a general conceptual drawback as indicated in the state of the art. In any event, electrical signals from the sensor are processed in terms of signal strength and duration to provide the microprocessor with information relating to the desired response. The sensing of a bending moment related to loading of the limb may be performed by means of a force sensing resistor mounted between, for example, relatively movable parts in the region of the knee joint, and a particularly preferred arrangement is to place the force sensing resistor between, on the one hand, one end of a lever arm the other end of which is pivoted on a knee chassis member forming part of the thigh part of the prosthesis, and, on the other hand, a resilient element inserted between the lever arm and the knee chassis to provide small amounts of flexion independently of the knee flexion control device. Other transducers may be used for producing an electrical signal, which is wholly or partly a function of the knee bending moment.
Blatchford clearly indicates a that a means is sought to determine a level of strain around the knee center to provide the controller with required information of the degree and duration of force, and feedback between resistive force output by the hydraulic controller, so that deviations from the desired levels can be used as input for changing the hydraulic controller's output by altering the state of stepper motor driven valves. The sensor in Blatchford is part of an integral electronic control system wherein the sensor provides feedback with regards to the system's response primarily to its own. Further, the specification of Blatchford provides a clear limitation of operability between 0-35° (GB page 2, 11.22-25), which is believed to arise, at least in part, from the nature of the reaction forces in hydraulic systems and the limitations placed on linear transducers: when the knee bends under resistance over a certain angle; not only does the mechanical leverage over the knee increase with progressive knee angle increasing signal strength in Blatchford's sensor, but also the shortening lever arm for operating piston with increasing knee flexion requires an ever higher hydraulic pressure inside to produce resistive bending moment to the knee. Certainly, using a system in accordance with such teaching, with a sensor in the knee region, will not allow operability of the hydraulic actuator in a meaningful way. This makes the relative placement of any sensor material and a better placement of sensor distinct if it can add new functionality. It is noted that the approach to placement of a sensor in an otherwise comprehensive specification appears to be quite random, nor are the ramifications mentioned.
In a specific example, in the event that a strong force is generated by a forward swing, inertia arising from a distal element will increase the mechanical force sensed by the sensor, and result in an over-estimate of the level of damping required, which would then reduce its resistance to balance the system response, making the response of the prosthesis too weak. Similarly, upon descent of a slope, a hip extension in a weight bearing condition will necessarily result that could cause an increase or decrease in the level of the force determined by the sensor, in dependence upon a momentary angle of knee flexion, and this would necessarily influence the information received by the electronic system. Certainly, using a system in accordance with such teaching, with a sensor positioned adjacent the knee (to minimize the above mentioned disadvantages in relation to stump generated forces affecting the sensor signals that are likely to register knee bending moments), will not allow efficient operability of the hydraulic actuator if sensor is placed remote from the knee center. In fact, if same anterior auxiliary axis and posterior sensor were place more proximal, then hip extension moment would cause sensor to be in tensile mode instead and not represent the moment about the knee joint it is meant to represent as a compressive force signal. The ability of reflex muscle actions to upset signal processing and knee behaviour is a disadvantage present invention seeks to avoid. This makes the relative placement of any sensor material and a better placement of sensor distinct if it can add new functionality. However, such an approach to placement of a sensor in an appears to be at odds with the other ramifications of the specific teaching mentioned in this document.
In another example, concerning the so-called Total Knee by Ossur, a linear damper is placed just proximal a pivot arrangement of a polycentric mechanism. In this knee, the true center in the swing phase the instantaneous center of rotation, which is proximal to the center of mass of the knee joint, as can be determine with reference to a 2007 paper by the “Ossur Academy” entitled Gait training and Prosthetic Knees” by E Kennedy and F Barnett. This paper shows that the linear damper of the Total Knee moves with the thigh element, and in that sense does not control its own inertia, but as a modular joint, its center of mass is distal to the effective knee axis at all times during swing and thigh element does not affect the state of operation of the damper.
The present invention seeks to provide a solution to the problems addressed above: problems associated with moment of inertia of the prosthetic knee device, the need to switch operational modes of the knee in response to ground reaction forces and body reflexes over a full range of knee flexion. The present invention seeks to provide an improved mechanically operated prosthesis for a short stumped femoral amputee, where ‘short stump’ is intended a length short enough to locate the center of mass of the knee joint and its main operable components proximal to the knee center. The present invention also seeks to provide an improved electrically operated prosthesis for a femoral amputee.
The present invention further seeks to provide a device that is commercially and clinically relevant, and that adequately deals with the problems apparent in respect of certain prior teachings of the art. The present invention further seeks to provide a response independent of a level of anterior-posterior force, as determined by some proponents of the art, with respect to a level of resistance as experienced by the prosthesis.